In recent years, significant improvements have been made in medical diagnostic instrumentation, particularly as related to the diagnosis of internal injuries and illnesses. In particular, the field of imaging internal organs and tumors in the diagnostic context has progressed dramatically in recent years.
A widely used modern imaging technique is generally referred to as magnetic resonance imaging (MRI). MRI utilizes the nuclear magnetic resonance (NMR) mechanism available from the nuclei of those atoms which are gyromagnetic (i.e., atomic nuclei that have nonzero angular momentum, or nonzero magnetic moment), the most commonly analyzed of which is hydrogen. MRI is accomplished by placing the subject to be imaged, for example a human, into a DC magnetic field of a magnitude ranging from 0.15 Tesla to 2.0 Tesla for polarizing the gyromagnetic nuclei of interest in the subject. A gradient, or sweep, coil is used to produce a gradient magnetic field within the subject, precisely defining the volume within the subject to be imaged at a particular time; modulation of the gradient field allows for sequential imaging of small volumes so that an image of a larger volume can be produced. As is well known in the art, the gradient field, in combination with the DC field, will define a volume in which the atoms will have a common resonant frequency different from atoms outside of the volume. The MRI apparatus also generally includes an oscillator coil that generates an oscillating magnetic field, oriented at an angle relative to the DC field; the frequency of the oscillating field is selected to match the resonant frequency of the atoms of interest in the selected volume. Frequencies of interest in modern MRI are in the radio frequency range. The MRI apparatus also includes a detecting coil in which a current can be induced by the nuclear magnetic dipoles in the volume being imaged.
In operation, as is well known, the magnetic dipole moments of those atoms in the volume which are both gyromagnetic and also resonant at the frequency of the oscillating field are rotated to some angle, for example 90.degree., from their polarized orientation by the resonant RF oscillation. Upon removal of the RF excitation, the voltage induced in the detecting coil is measured over time, with the rate of decay corresponding to the quantity of the atoms of interest in the volume being imaged. Sequencing of the selected volume by modulating the gradient field will thus result in an image of the subject which depends upon the density of the resonant nuclei (e.g., hydrogen) at the volumes therein. MRI according to this technique (and its variations) has been successful in the imaging of soft tissues which are transparent to X-rays, such as internal organs and the like.
Generation of large DC magnetic fields has been a particularly difficult problem in conventional MRI equipment. Firstly, imaging is facilitated by the generation of extremely large fields of on the order of at least 1.5 Tesla, for example. Such large magnetic fields require large coil currents or alternatively larger numbers of turns in the main DC magnet coil, each of which have necessitated the use of superconducting material and accompanying cryogenic systems in the magnets which are both relatively costly items. In addition, the size and weight of the MRI apparatus generally increases with the DC field strength of the magnet, as the weight and size of the magnet will increase with its DC field strength. Some conventional MRI magnets are sufficiently heavy (e.g., on the order of twenty tons) as to limit the installation of the MRI apparatus to a basement or ground floor laboratory.
In addition to these cost factors, an important factor in the design and construction of MRI magnets is the uniformity of the magnetic field within the "gap" into which the subject is to be placed, as the precision of the image depends upon the uniformity of the field at the subject location. It has been observed that magnetic field uniformity becomes increasingly difficult to achieve with increasing field strength.
In combination with these cost and uniformity factors, shielding of stray fields from the DC magnet also becomes of major concern, especially as the fields become very large, as the accuracy of conventional electronic instrumentation is adversely impacted by magnetic fields. Many purchasers of MRI equipment thus demand excellent shielding, as the size and cost of the laboratory into which the MRI apparatus is to be installed will depend not only upon the size of the apparatus itself, but also upon the distance that other instrumentation must be kept from the apparatus (or alternatively the external shielding measures required). Use of superconducting coils to provide such shielding can be effective, but of course dramatically increases the cost of the apparatus. Iron shielding, while avoiding the cost of additional superconducting material, results in much larger and heavier magnets, and also must be designed in a highly precise manner, especially for high fields, so that the iron does not adversely affect the uniformity of the field in the imaging volume.
By way of further background, a superferric shielded magnet and method for constructing the same is described in U.S. Pat. No. 4,783,628, issued Nov. 8, 1988, and in U.S. Pat. No. 4,822,772, both incorporated herein by this reference. A superconducting magnet particularly intended for MRI applications, shielded by the combination of a superferric shield and superconducting shielding coils, is described in copending application Ser. No. 715,552, filed Jun. 14, 1991, also incorporated herein by this reference. In each of these magnets, the bore is cylindrical, with the generated DC magnetic field in a direction parallel with the axis of the cylindrical bore.
By way of further background, MRI equipment utilizing a window frame opening into which the subject is placed are described in U.S. Pat. No. 4,791,370, issued Dec. 13, 1988, and U.S. Pat. No. 4,667,174, issued May 19, 1987. In these arrangements, an iron yoke is provided into which upper and lower U-shaped non-superconducting coil assemblies are provided for generating a vertical DC magnetic field within its volume. Gradient coils are inserted thereinto (as described in U.S. Pat. No. 4,791,370) for selecting the slice of the subject to be imaged.
The cylindrical and window frame magnets described hereinabove have each been widely used in MRI equipment intended for human subjects. However, access to the patient is limited in MRI equipment fabricated according to these shapes, particularly in the case of cardiac patients to whom intravenous tubes and cardiac monitoring equipment must remain connected during the MRI procedure. As MRI is a particularly useful diagnostic tool for cardiac disorders, it is therefore highly desirable to provide an MRI apparatus which can easily receive a human subject and yet provide external access to the subject. In addition, MRI equipment using the cylindrical and window frame magnet designs can result in significant patient anxiety, as these shapes require the head of the patient to be enclosed within the magnet during the MRI of many regions of the body.
By way of still further background, C-shaped magnets for generating DC fields are well known in the field of particle acceleration. A well known example of such magnets are those used in cyclotrons to maintain a centripetal acceleration on the accelerated charged particle(s) traveling between the accelerating dee electrodes. Various pole face designs have been utilized in cyclotron magnets in order to properly focus the path traveled by the particle, especially as it is accelerated by the dee electrodes.
By way of further background, a C-shaped magnet for generating a magnetic field intended for use in NMR equipment is described in international publication WO 90/04258, published Apr. 19, 1990, particularly relative to FIG. 3 therein. In the C-shaped magnet described in this document, an iron yoke is provided to serve as a flux guide. An array of modular solenoids are used to generate the field across the gap; the publication also discloses that the modules can be powered to have different energy densities from one another to maximize the efficiency of the array of modules.
It is an object of the present invention to provide a high field DC magnet intended for MRI equipment which provides easy patient access and which does not require the patient to be enclosed within the magnet during the MRI procedure.
As will be described hereinbelow, a C-shaped magnet according to the invention provides excellent patient access, and can also perform MRI of many parts of the human body without enclosing the head area, thus reducing patient anxiety. However, the C-shape of the magnet theoretically provides poor uniformity of return flux to the field generating coils at the gap area. This poor uniformity results from the natural crowding of magnetic flux near the inner surface of a curved flux path (i.e., at the inner surfaces of the top and bottom of the "C"). As high uniformity of field within the imaging volume (e.g., a sphere of 20 cm diameter) is essential in the MRI context for precise imaging, the improved patient access provided by a C-shaped magnet adversely impacts the uniformity of the field in the gap.
It is therefore a further object of the invention to provide such a magnet which has a high degree of field uniformity in the gap.
It is a further object of the invention to provide such a magnet which provides such high field uniformity by correcting for flux crowding in the curved return path.
It is a further object of the invention to provide such a magnet which has a high degree of shielding and low fringe fields.
It is a further object of the invention to provide such a magnet which can be fabricated at relatively low cost.
Other objects and advantages of the present invention will be apparent to those of ordinary skill in the art having reference to the following specification in combination with the drawings.